The human retina is sensitive to alterations in blood flow provided by the retinal vascular system. Such alterations in blood flow may: (a) occur as a result of a specific ocular disease such as blood vessel occlusion or (b) be associated with a systemic disease such as diabetes or systemic hypertension, which alterations in blood flow often lead to blindness. In light of this, an efficient, reproducible method for measuring blood flow in the retina is important for use in diagnosis and treatment of retinal vascular diseases.
A number of apparatus are known in the art for measuring/estimating blood flow in the retina, however all suffer from one or more drawbacks. One such prior art apparatus is an ultrasound imaging system which measures a Doppler-shift in a reflected wave. The ultrasound imaging system uses a first technique to provide output to a user in a format known as Spectral Doppler flow mapping and a second technique to provide output to the user in a format known as Color Doppler flow mapping, which first and second techniques differ mainly in the manner in which data is presented to the user. The ultrasound imaging system suffers from a drawback in that neither the first nor the second technique can provide accurate three-dimensional localization of a signal, nor can these techniques account for the direction of blood flow relative to a direction of sound wave propagation (known in the art as "transducer alignment"). The inherent uncertainty in blood flow direction, referred to as transducer alignment error, can have large effects on the measured velocity. Furthermore, filter settings in the ultrasound apparatus can also have large effects on the absolute measured velocities. Filter induced errors, combined with transducer alignment errors, have the consequence that ultrasound Doppler techniques can provide only an averaged estimate of retinal blood flow.
Another such prior art apparatus is a Confocal Scanning Laser Ophthalmoscope ("CSLO") which can provide accurate three-dimensional images of the anterior eye, including the retina. Blood flow measurements in the retina have been made using the CSLO instrument by means of an indirect technique wherein fluorescent dyes are injected into the blood stream and video images of vascular structures under laser illumination are provided to show the spreading fluorescence. An indication of blood flow is provided by the strength of the time-dependent fluorescence. This technique suffers from a drawback in that it requires extensive post-processing of video images by a trained operator to extract gray scales and to correlate the gray scales with blood flow, see an article entitled "Retinal Circulation Time Determination using the SLO--Image Processing Techniques" by P. G. Rehkopf, J. W. Warnicki, L. J. Mandarino, T. R. Friberg, and D. N. Finegold, paper presented at the First International Symposium on Scanning laser Ophthalmoscopy and Tomography, University Eye Hospital, Munich, Germany, Jul. 7-8, 1989. A further drawback of this technique is that it requires injection of dyes into the blood stream.
Still another such prior art apparatus is a Laser Doppler Velocimeter ("LDV") which measures blood flow velocity in the retina by attaching a laser/detector system to a standard fundus camera, see an article entitled "Near-IR Retinal Laser Doppler Velocimetry and Flowmetry: New Delivery and Detection Techniques" by B. L. Petrig and C. E. Riva, Applied Optics 30, 1 Jun. 1991, pp. 2073-2078. The LDV technique can provide localized (on a 2-D plane), absolute measurement of blood flow velocity in major retinal vessels (&gt;50 .mu.m diameter) in a non-invasive manner. However, the standard LDV technique suffers from several drawbacks. First, the LDV technique suffers a drawback in that it cannot provide 3-D localization by itself. Second, the LDV technique suffers a drawback in that it includes simultaneous measurement of all blood velocities within a blood vessel (as is well known, blood flows fastest in the center of the blood vessel and slowest at the wall). The presence of all blood velocities within a detection signal: (a) complicates the analysis and, (b) when short sampling times are used to avoid multiple scattering artifacts, makes interpretation of the detection signal difficult. This is illustrated in FIG. 1A where, for the LDV technique, the coherence length 210 of laser beam 220 is much greater than the diameter of blood vessel 200. Third, the LDV technique suffers a drawback in that it includes multiple scattering artifacts. Multiple scattering effects are disadvantageous because multiple scattering tends to dominate the reflected signal over long sampling times, thereby making velocity determination difficult. As a result, one is forced to use short sampling times. This is illustrated in FIGS. 1B and 1C. FIG. 1B shows, in graphical form, a frequency spectrum of a measured LDV signal for an idealized blood sample, i.e., a very dilute blood sample in which only single scattering need be considered. For this idealized case of single scattering, the measured LDV signal frequency spectrum is constant with respect to frequency up to a maximum frequency, f.sub.max, where it abruptly drops to the shot noise limit. This behavior is caused by the parabolic red blood cell velocity profile found in blood vessels. The parabolic velocity profile causes each velocity increment (corresponding to a Doppler frequency increment) to contribute an equal amount to the measured LDV signal strength and this produces a flat frequency spectrum. Beyond f.sub.max there is no further signal, except for the noise terms. In contrast to the idealized case of single scattering shown in FIG. 1B, FIG. 1C shows a representative spectrum from actual blood vessels wherein multiple scattering forces the use of very short sample times. The use of short sample times, in turn, makes exact determination of f.sub.max difficult because the reflected signal is naturally stochastic and, as a result, short sampling times increase the noise. However, long sample times allow the signal beyond f.sub.max to build up (multiple scattering can cause frequency shifts larger than f.sub.max). Therefore, determining the cut-off frequency f.sub.max is possible only for short sampling times. Fourth, the LDV technique suffers a drawback in that it relies on reflection from the blood vessel wall as a "local oscillator" (the blood vessel wall provides a strong, non-Doppler-shifted reference beam). Interference between the local oscillator beam and the Doppler-shifted signal causes a low-frequency beat signal in the detector from which the velocity can be extracted. However, the beat frequencies are typically a few kilohertz and are therefore sensitive to low frequency 1/f noise which is inherent in the LDV apparatus.
In light of the above, there is a need in the art for an apparatus for providing non-invasive, blood flow velocity measurement which overcomes the above-described problems.